Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc.
There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a bodily region of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis.
In traditional PET imaging, a patient is injected with a radioactive substance with a short decay time. As the substance undergoes positron emission decay, it emits positrons which, when they collide with electrons in the patient's tissue give off two gamma rays. The gamma rays emerge from the patient's body simultaneously at substantially opposite directions. A number of these gamma ray pairs eventually reach scintillation devices positioned in opposing locations around the patient. The scintillation devices often are configured as a ring of scintillation devices surrounding the patient. When each of the gamma rays of the pair interact with the scintillation devices, a burst of light is emitted and detected. The light is usually transmitted through a lightguide to a photodetector where the light photons are converted to an electrical signal. The electrical signals produced by the photodetector are then interpreted by a processor and accumulated, from which an image of the region of interest may be reconstructed.
Gamma-ray signals from a scintillator such as LSO (lutetium oxyorthosilicate) and BGO (bismuth germanate) have an intrinsic shape, the signals have a fast rising edge following a slow falling edge. The signals can be estimated as a function of:
            V      0        ⁡          (      t      )        ≈            A      1        ×          m      0        ×          (                                    1                          τ              1                                ×                      ⅇ                                          -                t                            /                              τ                1                                                    -                              1                          τ              0                                ×                      ⅇ                                          -                t                            /                              τ                0                                                        )      where the decay time-constant τ0 is determined by the scintillation crystal, and time-constant τ1 is mainly determined by the characteristics of the photosensor, the open-loop gain of the first amplifier in the front-end electronics, and the input capacitance (A1 and m0 are factors associated with the number of emitted photons from the crystal in response to excitation by a gamma ray and conversion to a voltage). When τ0>>τ1 (which is the case for LSO and BGO crystals), τ1 dominates the rising edge of the signal pulse, and τ0 dominates the pulse falling edge. The Laplace transfer-function of the above equation is:
      H    ⁡          (      s      )        =                    A        1                    τ        1              ×                  s                              (                          s              +                              1                                  τ                  0                                                      )                    ×                      (                          s              +                              1                                  τ                  1                                                      )                              .      
The falling edge of the scintillation signal is a first-order exponential decay function, so the shape of the signal is always unipolar; it is either positive or negative depending on the electronic readout circuits used.
“Multiplexing” as used in NM electronics refers to encoding of combinations of signals from different photodetectors or photodetector segments to determine the spatial location of a gamma event in the scintillator. See U.S. Pat. No. 3,011,057 to Anger, incorporated herein by reference in its entirety, in particular FIG. 2. Resistor network based multiplexing methods also have been invented and implemented in “Position-Sensitive PMT” (PS-PMT) and “Multi-Channel PMT” (MC-PMT) PET/PEM (Positron Emission Mammography) systems. These resistor networks process signals only from current sources, such as PS-PMT and MC-PMT anode outputs. Thus, such resistor-based networks cannot be implemented in voltage source detectors. For example, an Avalanche-Photo-Diode (APD) PET detector must have a charge-sensitive amplifier to amplify the APD output at the first stage. Naturally, subsequent amplifier outputs are voltage sources rather than current sources.
More importantly, the resistor network schemes have limited signal dynamic range. The “multiplexing” detector blocks will share one position histogram image. So histogram images and position lookup tables are actually not multiplexed event though the signal channels are. This mapping (or multiplexing) design could cause poorer crystal identification ability, ultimately leading to potential degradation of PET image resolution.
In a combination PET system such as MR/PET, the PET main electronics cannot be close to the MR scanner. Practically, they need to be located outside, in a MR RF-shielded room. In this case, longer signal transmission cables are needed to connect between the PET detectors and the main electronics. Cable shielding and grounding potential could become an issue.
Therefore, it is desired to have multiplexing design that can be implemented for both current and voltage detector sources, and wherein each block can have its own position histogram image and lookup table and the scintillation crystal can be better identified, so the nearby crystal elements have less crosstalk problems.